This invention relates to the field of medical devices, and more particularly to guiding means such as guidewires for advancing catheters within body lumens in procedures such as percutaneous transluminal coronary angioplasty (PTCA).
In typical PTCA procedures a guiding catheter having a preformed distal tip is percutaneously introduced into the cardiovascular system of a patient in a conventional Seldiger technique and advanced therein until the distal tip of the guiding catheter is seated in the ostium of a desired coronary artery. A guidewire is positioned within an inner lumen of a dilatation catheter and then both are advanced through the guiding catheter to the distal end thereof. The guidewire is first advanced out of the distal end of the guiding catheter into the patient""s coronary vasculature until the distal end of the guidewire crosses a lesion to be dilated, then the dilatation catheter having an inflatable balloon on the distal portion thereof is advanced into the patient""s coronary anatomy over the previously introduced guidewire until the balloon of the dilatation catheter is properly positioned across the lesion. Once in position across the lesion, the balloon is inflated to a predetermined size with radiopaque liquid at relatively high pressures (e.g. greater than 4 atmospheres) to compress the arteriosclerotic plaque of the lesion against the inside of the artery wall and to otherwise expand the inner lumen of the artery. The balloon is then deflated so that blood flow is resumed through the dilated artery and the dilatation catheter can be removed therefrom.
Conventional guidewires for angioplasty and other vascular procedures usually comprise an elongated core member with one or more tapered sections near the distal end thereof and a flexible body such as a helical coil disposed about the distal portion of the core member. A shapable member, which may be the distal extremity of the core member or a separate shaping ribbon which is secured to the distal extremity of the core member extends through the flexible body and is secured to a rounded plug at the distal end of the flexible body. Torquing means are provided on the proximal end of the core member to rotate, and thereby steer, the guidewire while it is being advanced through a patient""s vascular system.
Further details of dilatation catheters, guidewires, and devices associated therewith for angioplasty procedures can be found in U.S. Pat. No. 4,323,071 (Simpson et al.); U.S. Pat. No. 4,439,185 (Lundquist); U.S. Pat. No. 4,516,972 (Samson); U.S. Pat. No. 4,538,622 (Samson et al.); U.S. Pat. No. 4,554,929 (Samson et al.); U.S. Pat. No. 4,616,652 (Simpson); and U.S. Pat. No. 4,638,805 (Powell).
Steerable dilatation catheters with fixed, built-in guiding members, such as described in U.S. Pat. No. 4,582,181 (now Re 33,166) are frequently used because they have lower deflated profiles than conventional over-the-wire dilatation catheters and a lower profile allows the catheter to cross tighter lesions and to be advanced much deeper into a patient""s coronary anatomy.
A major requirement for guidewires and other guiding members, whether they be solid wire or tubular members, is that they have sufficient columnar strength to be pushed through a patient""s vascular system or other body lumen without kinking. However, they must also be flexible enough to avoid damaging the blood vessel or other body lumen through which they are advanced. Efforts have been made to improve both the strength and flexibility of guidewires to make them more suitable for their intended uses, but these two properties are for the most part diametrically opposed to one another in that an increase in one usually involves a decrease in the other.
The prior art makes reference to the use of alloys such as NITINOL, which is an acronym for Ni-TiNaval Ordnance Laboratory. These alloys have shape memory and/or superelastic characteristics and may be used in medical devices which are designed to be inserted into a patient""s body. The shape memory characteristics allow the devices to be deformed to facilitate their insertion into a body lumen or cavity and then be heated within the body so that the device returns to its original shape. Superelastic characteristics on the other hand generally allow the metal to be deformed and restrained in the deformed condition to facilitate the insertion of the medical device containing the metal into a patient""s body, with such deformation causing the phase transformation. Once within the body lumen the restraint on the superelastic member can be removed, thereby reducing the stress therein so that the superelastic member can return to its original undeformed shape by the transformation back to the original phase.
Alloys having shape memory/superelastic characteristics generally have at least two phases. These phases are a martensite phase, which has a relatively low tensile strength and which is stable at relatively low temperatures, and an austenite phase, which has a relatively high tensile strength and which is stable at temperatures higher than the martensite phase.
Shape memory characteristics are imparted to the alloy by heating the metal at a temperature above which the transformation from the martensite phase to the austenite phase is complete, i.e. a temperature above which the austenite phase is stable. The shape of the metal during this heat treatment is the shape xe2x80x9crememberedxe2x80x9d. The heat treated metal is cooled to a temperature at which the martensite phase is stable, causing the austenite phase to transform to the martensite phase. The metal in the martensite phase is then plastically deformed, e.g. to facilitate the entry thereof into a patient""s body. Subsequent heating of the deformed martensite phase to a temperature above the martensite to austenite transformation temperature causes the deformed martensite phase to transform to the austenite phase and during this phase transformation the metal reverts back to its original shape.
The prior methods of using the shape memory characteristics of these alloys in medical devices intended to be placed within a patient""s body presented operational difficulties. For example, with shape memory alloys having a stable martensite temperature below body temperature, it was frequently difficult to maintain the temperature of the medical device containing such an alloy sufficiently below body temperature to prevent the transformation of the martensite phase to the austenite phase when the device was being inserted into a patient""s body. With intravascular devices formed of shape memory alloys having martensite-to-austenite transformation temperatures well above body temperature, the devices could be introduced into a patient""s body with little or no problem, but they had to be heated to the martensite-to-austenite transformation temperature which was frequently high enough to cause tissue damage and very high levels of pain.
When stress is applied to a specimen of a metal such as NITINOL exhibiting superelastic characteristics at a temperature above which the austenite is stable (i.e. the temperature at which the transformation of martensite phase to the austenite phase is complete), the specimen deforms elastically until it reaches a particular stress level where the alloy then undergoes a stress-induced phase transformation from the austenite phase to the martensite phase. As the phase transformation proceeds, the alloy undergoes significant increases in strain but with little or no corresponding increases in stress. The strain increases while the stress remains essentially constant until the transformation of the austenite phase to the martensite phase is complete. Thereafter, further increase in stress are necessary to cause further deformation. The martensitic metal first yields elastically upon the application of additional stress and then plastically with permanent residual deformation.
If the load on the specimen is removed before any permanent deformation has occurred, the martensitic specimen will elastically recover and transform back to the austenite phase. The reduction in stress first causes a decrease in strain. As stress reduction reaches the level at which the martensite phase transforms back into the austenite phase, the stress level in the specimen will remain essentially constant (but substantially less than the constant stress level at which the austenite transforms to the martensite) until the transformation back to the austenite phase is complete, i.e. there is significant recovery in strain with only negligible corresponding stress reduction. After the transformation back to austenite is complete, further stress reduction results in elastic strain reduction. This ability to incur significant strain at relatively constant stress upon the application of a load and to recover from the deformation upon the removal of the load is commonly referred to as superelasticity or pseudoelasticity.
The prior art makes reference to the use of metal alloys having superelastic characteristics in medical devices which are intended to be inserted or otherwise used within a patient""s body. See for example, U.S. Pat. No. 4,665,905 (Jervis) and U.S. Pat. No. 4,925,445 (Sakamoto et al.).
The Sakamoto et al. patent discloses the use of a nickel-titanium superelastic alloy in an intravascular guidewire which could be processed to develop relatively high yield strength levels. However, at the relatively high yield stress levels which cause the austenite-to-martensite phase transformation characteristic of the material, it did not have a very extensive stress-induced strain range in which the austenite transforms to martensite at relative constant stress. As a result, frequently as the guidewire was being advanced through a patient""s tortuous vascular system, it would be stressed beyond the superelastic region, i.e. develop a permanent set or even kink which can result in tissue damage. This permanent deformation would generally require the removal of the guidewire and the replacement thereof with another.
Products of the Jervis patent on the other hand had extensive strain ranges, i.e. 2 to 8% strain, but the relatively constant stress level at which the austenite transformed to martensite was very low, e.g. 50 ksi.
What has been needed and heretofore unavailable is an elongated solid or tubular body for intravascular devices, such as guide wires or guiding members, which have at least a portion thereof exhibiting superelastic characteristics including an extended strain region over a relatively constant high stress level which effects the austenite transformation to martensite and still provide a one-to-one torque response. The present invention satisfies these and other needs.
The present invention is directed to an improved superelastic body which is suitable for intraluminal devices, such as guidewires or guiding members, wherein superelastic characteristics result from the stress-induced transformation of austenite to martensite.
The superelastic alloy body of the invention has a stable austenite phase which will transform to martensite phase upon the application of stress and will exhibit a recoverable strain of at least about 4% upon the stress induced transformation of the austenite phase to the martensite phase. The formation of the alloy body includes a final cold working about 10 to about 75% and then a final memory imparting heat treatment at a temperature of about 450xc2x0 to about 600xc2x0 C., preferably 475xc2x0 to about 550xc2x0 C. The cold worked, heat treated product exhibits a stress-induced phase transformation at temperatures below about 45xc2x0 C. at a relatively high stress level, e.g. above about 70 ksi, (483 Mpa) preferably above about 90 ksi (620 Mpa) for solid products and about 40 ksi (276 Mpa) for hollow tubular products. The superelastic product exhibits at recoverable strain of at least 4% upon completion of the stress-induced transformation of the austenite phase to the martensite phase. The onset of the stress induced phase change from austenite to martensite, preferably begins when the specimen has been strained about 2% and extends to a strain level of about 8% at the completion of the phase change. The stress and strain referred to herein is measured by tensile testing. The stress-strain relationship determined by applying a bending moment to a cantilevered specimen is slightly different from the relationship determined by tensile testing because the stresses which occur in the specimen during bending are not as uniform as they are in tensile testing. The stress change during the phase transformation is much less than the stress either before or after the stress-induced transformation. In some instances the stress level during the phase change is almost constant.
The elongated portion of the guiding member having superelastic properties is preferably formed from an alloy consisting essentially of about 30 to about 52% titanium, about 38% to about 52% nickel and additional alloying elements in amount up to 20% for copper and up to about 10% in the case of other alloying elements. The additional alloying elements may be selected from the group consisting of up to 3% each of iron, cobalt, chromium, platinum, palladium, zirconium, hafnium and niobium and up to about 10% vanadium. At nickel levels above 52% the alloy becomes too brittle to fabricate by cold working. Metallurgically the alloy consist essentially of a predominant quantity of a NiTi intermetallic compound and small quantities of other constituents. Additionally, when nickel is in excess Ni3Ti is formed, whereas when titanium is in excess Ti2Ni is formed. As used herein, all references to percent alloy compositions are atomic percent unless otherwise noted.
To form the elongated superelastic portion of the guiding member, elongated solid rod or tubular stock of the preferred alloy material is first thermomechanically processed through a series of cold working, e.g. drawing and inter-annealing at temperatures between about 600xc2x0 to about 800xc2x0 C. for about 5 to about 30 minutes and is then given a final cold working, e.g. drawing, to effect a final size reduction of about 10% up to about 75% in the transverse cross section thereof. For solid products the final cold work is preferably about 30 to about 70% and for hollow tubular products the final cold work is preferably about 10% to about 40%. As used herein % cold work is the size reduction of the transverse dimension of the work piece effected by the cold working. After the final cold working, the material is given a heat treatment at a temperature of about 450xc2x0 to about 600xc2x0 C., preferably about 475xc2x0 to about 550xc2x0 C., for about 0.5 to about 60 minutes to generate the superelastic properties. To impart a straight memory, the cold worked material may be subjected to a longitudinal stress equal to about 5% to about 50%, preferably about 10% to about 30%, of the yield stress of the material (as measured at room temperature) during a heat treatment of about 450xc2x0 to about 600xc2x0 C. This thermomechanical processing imparts a straight xe2x80x9cmemoryxe2x80x9d to the superelastic portion and provides a relatively uniform residual stress in the material. Preferably, the final cold worked product is subjected to mechanically straightening between the final cold working and heat treating steps to provide the wire or tubular product with substantially improved, one-to-one torque response, i.e. it is substantially whip free. The alloy composition and thermal treatment are selected to provide an austenite finish transformation temperature generally about xe2x88x9220xc2x0 to about 40xc2x0 C. and usually less than body temperature (approx. 37xc2x0 C.). To obtain more consistent final properties, it is preferred to fully anneal the solid rod or tubular stock prior to cold working so that the material will always have the same metallurgical structure at the start of the cold working. The pre-annealing also ensures adequate ductility for cold working. It will be appreciated by those skilled in the art that the alloy can be cold worked in a variety of ways other than drawing, such as rolling or swaging. The constant stress levels for tubular products have been found to be slightly lower than the constant stress levels for solid products due to the inability to cold work the tubular products to the extent the solid products can be cold worked. For example, solid superelastic wire material of the invention can have a relatively constant stress level above about 70 ksi, usually above about 90 ksi, whereas, hollow superelastic tubing material of the invention can have a relatively constant stress level above about 50 ksi (345 Mpa) usually above about 60 ksi (414 Mpa). The ultimate tensile strength of both forms of the material is well above 200 ksi (1380 Mpa) with an ultimate elongation at failure of about 15%.
The elongated body of the invention exhibits a stress-induced austenite-to-martensite phase transformation over a broad region of strain at very high, relatively constant stress levels. As a result a guiding member formed of this material is very flexible, it can be advanced through very tortuous passageways such as a patient""s coronary vasculature with little risk that the superelastic portion of the guiding member will develop a permanent set and at the same time it will effectively transmit the torque applied thereto without causing the guiding member to whip.
These and other advantages of the invention will become more apparent from the following detailed description thereof when taken in conjunction with the following exemplary drawings.